There are a number of medical treatments, such as ultrafiltration, apheresis and dialysis, that require blood to be temporarily withdrawn from a patient and returned to the body shortly thereafter. While the blood is temporarily outside of the body, it flows through an “extracorporeal blood circuit” of tubes, filters, pumps and/or other medical components. In some treatments, the blood flow is propelled by the patient's blood pressure and gravity, and no artificial pump is required. In other treatments, blood pumps in the extracorporeal circuit provide additional force to move the blood through the circuit and to control the flow rate of blood through the circuit. These pumps may be peristaltic or roller pumps, which are easy to sterilize, are known to cause minimal clotting and damage to the blood cells, and are inexpensive and reliable.
Brushed and brushless DC motors are commonly used to rotate peristaltic pumps. A motor controller regulates the rotational speed of blood pumps. The speed of a pump, expressed as rotations per minute (RPM), regulates the flow rate of the blood through the circuit. Each revolution of the pump moves a known volume of blood through the circuit. Thus, the blood flow rate through the circuit can be easily derived from the pump speed. Accordingly, the pump speed provides a relatively accurate indicator for the volume flow of blood through an extracorporeal circuit.
Existing pump controllers may be as simple as a potentiometer that regulates the voltage to the pump DC motor. The pump speed is proportional to the voltage applied to the pump motor. By increasing the voltage, the speed of the pump increases and, similarly, the blood flow increases through the extracorporeal circuit. More sophisticated existing pump controllers, such as used in current dialysis machines, include a microprocessor that executes a software/firmware program to regulate the pump speed and, thus, blood flow in accordance with pump/flow settings entered by an operator. In these controllers, the microprocessor receives input commands from an operator who selects a desired blood flow using a user interface on the controller housing. The microprocessor determines the pump motor speed needed to provide the selected blood flow rate, and then issues commands to the pump motor to run at the proper speed.
To improve the accuracy and reliability of the blood flow through an extracorporeal circuit, existing pump controller microprocessors receive feedback signals from, for example, tachometers or optical encoders that sense the actual speed of the pump motor. Similarly, feedback signals have been provided by ultrasonic flow probes that measure the actual flow of blood in the extracorporeal circuit. By comparing the desired pump speed or flow rate to the measured pump speed or flow rate, the microprocessor can properly adjust the pump speed to correct for any difference between the desired and actual speed or rate. In addition, a calculated or measured flow value (actual flow rate) may be displayed on the pump console for viewing by the operator for a visual comparison with the desired flow setting.
The microprocessor controllers for blood pumps have, in the past, relied on both open and closed loop control of the motor speed. The open loop control normally consists of a constant feed forward voltage (based on the back EMF constant of the motor). The closed loop control systems use velocity feedback in the form of a tachometer, encoder or resolver to maintain constant pump flow. A constant flow control loop allows a user to set the blood flow rate, and the controller regulates the pump speed to maintain a constant blood flow, unless a malfunction occurs that would require the pump to be shut-down to protect the patient. The open loop control systems have the disadvantage that an increase or decrease in torque or motor resistance due to temperature will result in some variation in pump flow. This variation will generally be small when the motor is geared. Torque variations are not an issue for closed loop control systems because they use the motor velocity as feedback and adjust the supply current or voltage to the motor in order to maintain constant velocity.
Existing blood pump controllers include various alarms and interlocks that are set by a nurse or a medical technician (collectively referred to as the operator), and are intended to protect the patient. In a typical dialysis apparatus, the blood withdrawal and blood return pressures are measured in real time, so that sudden pressure changes are quickly detected. Sudden pressure changes in the blood circuit are treated as indicating an occlusion or a disconnect in the circuit. The detection of a sudden pressure change causes the controller to stop the pump and cease withdrawal of blood. The nurse or operator sets the alarm limits for the real time pressure measurements well beyond the expected normal operating pressure for the selected blood flow, but within a safe pressure operating range.
Examples of existing blood pump controllers are disclosed in U.S. Pat. No. 5,536,237 ('237 patent) and U.S. Pat. No. 4,657,529 ('529 patent). The controllers disclosed in these patents purport to optimize the rate of blood flow through a blood circuit based on a pressure vs. flow rate control curve. However, these patents do not teach controlling a blood pump based on control curves for both withdrawal and infusion pressures, and do not suggest reversing blood flow to relieve an occlusion. The authors of the '237 and '529 patents acknowledged that during blood withdrawal, treatment is often interrupted if the vein is collapsed. They further acknowledged that as the result of such collapse the needle could be drawn into the blood vessel wall that makes the recovery difficult. The remedy proposed by these authors is to always withdraw blood at a rate that prevents the collapse of the vein by applying a complex system of identification of the vein capacity prior to treatment. The '237 patent specifically addresses the difficulty of generating ad-hoc pressure flow relationships for individual patients with low venous flow capacity. However, the experience of the present applicants is that the withdrawal properties of venous access in many patients are prone to frequent and often abrupt changes during treatment. Accordingly, the approach advanced in the '237 and '529 patents of attempting to always avoid vein collapse will fail when a vein collapse condition occurs and does not provide any remedy for vein collapse (other than to terminate treatment and call a nurse or doctor).
Pressure conditions in a blood circuit often change because of blood viscosity changes (that in turn affect flow resistance), and because of small blood clots that form where blood stagnates on the surface of tubes and cannulae. These small clots partially occlude the blood circuit, but do not totally obstruct blood flow through the circuit. These clot restrictions increase the flow resistance and, thus, increase the magnitude of pressure in the circuit. Accordingly, an assumption that the flow resistance and pressure are constants in a blood circuit may not be valid. Gravity is another source of pressure change in a blood circuit. Gravity affects the pressure in a blood circuit. The pressure in the circuit will change due to gravity if the patient moves such that the height chances occur with respect to the vertical position of the circuit entry and exit points on the patient's arms relative to the pressure sensors. The pressure sensors in the blood circuit may detect a change that is indicative of a change in the patient's position, rather than a clot in the circuit. For example, if the patient sits up, pressure sensors will detect pressure changes of the blood in the circuit.
If the gravity induced pressure changes are sufficiently large, prior pump controllers tended to activate an alarm and shut down the blood pump. The operator then has to respond to the alarm, analyze the situation and remedy the malfunction or change the alarm limits if the flow path conditions have changed. In more advanced systems the operator sets the alarm window size (for example, plus or minus 50 mmHg about a mean point), and the machine will automatically adjust the mean point in proportion to the flow setting change. If the measured pressure exceeds the pressure range set by the window, the blood flow is automatically stopped. However, existing systems do not adjust the pump speed in response to changes in flow resistance, but rather, shut down if the resistance (and hence fluid pressure) become excessive.